Temperature profile mapping and guided thermotherapy

ABSTRACT

Techniques, apparatus and systems that use an optical probe head to deliver light to a target and to collect light from the target for imaging and monitoring a target while a separate radiation is applied to treat the target.

PRIORITY CLAIM

This patent document claims the priority of U.S. Provisional ApplicationNo. 61/032,853 entitled “Temperature Profile Mapping with Tomophase OCTApparatus” and filed on Feb. 29, 2008, the entire disclosure of which isincorporated by reference as part of the disclosure of this document.

BACKGROUND

This application relates to techniques, apparatus and systems that useoptical waveguides to deliver light to a targeted area for opticaldetection of tissues, organs and other objects in medical, biologicaland other applications.

Light can be guided through a light pipe or optical waveguide such asoptic fiber to a target to obtain optical images, optical measurementsand other operations of the target. The optical waveguide such as opticfiber can be used to reach the target at a location that is otherwisedifficult to reach or requires some preparatory procedures to make thetarget more accessible. For example, the tissue of an internal organ ofa patient may be made available for a medical examination or therapyprocedure through a natural orifice or an incision to expose theinternal organ. Such a procedure may be performed by delivering probelight to the tissue via an endoscope instrument or catheter to reduce orminimize the degree of invasiveness of a diagnostic or therapeuticprocedure. At the distal end of the instrument, light is pointed tocertain direction or steered to interact with an area or a slice oftissue of interest.

Various devices and techniques based on optical coherence domainreflectometry (OCDR) may be used for non-invasive optical probing ofvarious substances, including but not limited to skins, body tissues andorgans of humans and animals, to provide tomographic measurements ofthese substances. In many OCDR systems, the light from a light source issplit into a sampling beam and a reference beam which propagate in twoseparate optical paths, respectively. The light source may be partiallycoherent source. The sampling beam is directed along its own opticalpath to impinge on the substances under study, or sample, while thereference beam is directed in a separate path towards a referencesurface. The beams reflected from the sample and from the referencesurface are then brought to overlap with each other to opticallyinterfere. Because of the wavelength-dependent phase delay, the opticalinterference results in no observable interference fringes unless thetwo optical path lengths of the sampling and reference beams aresimilar. This provides a physical mechanism for ranging. A beam splittermay be used to split the light from the light source and to combine thereflected sampling beam and the reflected reference beam for detectionat an optical detector. The application of OCDR in medical diagnoses incertain optical configurations has come to be known as “opticalcoherence tomography” (OCT).

SUMMARY

This application includes implementations and examples of techniques,apparatus and systems that use an optical probe head in an endoscopedevice to optically measure a target for various applications, includingtechniques, apparatus and systems that use an optical probe head todeliver light to a target and to collect light from the target forimaging and monitoring a target while a separate radiation is applied totreat the target.

In one aspect, an implementation of an endoscope device includes anoptical catheter comprising an optical fiber to guide an optical imagingbeam and an optical probe head, located at a distal end of the fiber todirect the optical imaging beam from the fiber to a target and toreceive light returned from the target under illumination of the opticalimaging beam; a detection module to process the light returned from thetarget under illumination of the optical imaging beam to measure atemperature at a location illuminated by the optical imaging beam; anenergy source that produces energy that is applied to the target toraise a temperature at a location of the target where the energy isapplied; and a control mechanism that controls the energy source to seta power level of the energy produced by the energy source based on themeasured temperature from the detection module.

In another aspect, an implementation of a method for guiding theapplication of a thermotherapeutic radiation to a target tissue uses anoptical catheter comprising an optical fiber to guide an optical imagingbeam to direct the optical imaging beam from the fiber to a target andto receive light returned from the target under illumination of theoptical imaging beam, processes the light returned from the targettissue under illumination of the optical imaging beam to measure atemperature at a location illuminated by the optical imaging beam;applies a thermotherapeutic energy to the target tissue to raise atemperature at a location of the target tissue where the energy isapplied; and controls an amount of the thermotherapeutic energy appliedto the target tissue, based on the measured temperature by using theoptical imaging beam, to control the temperature at the target tissuebetween a low limit above which a thermotherapeutic effect is presentand a high limit above which a damage to the target tissue occurs.

In yet another aspect, an implementation of an endoscope device forproviding guided thermotherapy includes an endoscope tube comprising ahollow working channel and an optical catheter comprising an opticalfiber located inside the hollow working channel to guide an opticalprobe beam and, an optical probe head, located at a distal end of thefiber, to reflect a first portion of the optical probe beam back to thefiber and to direct a section portion of the optical probe beam to atarget tissue as an optical imaging beam. The optical probe headreceives light returned from the target tissue under illumination of theoptical imaging beam to overlap the light returned from the targettissue with the first portion to co-propagate in the fiber away from theoptical probe head. This device includes an optical delay device coupledto the fiber to receive the first portion and the light returned fromthe target tissue to produce a variable relative phase delay between thefirst portion and the light returned from the target tissue; an opticaldetector that detects the light of the first portion and the lightreturned from the target tissue from the optical delay device; aprocessing unit to receive output from the optical detector and toextract a temperature at a location illuminated by the optical imagingbeam from information of the target carried by the light returned fromthe target tissue; an RF applicator engaged to the endoscope tubing andnear the optical probe head to apply RF energy to the target tissue toraise a temperature at a location of the target where the RF energy isapplied; and a control mechanism that controls an amount of the RFenergy to be applied by the RF applicator to the target tissue based onthe measured temperature.

These and other aspects of various techniques, apparatus and systems aredescribed in detail in the drawings, the description and the claims.

BRIEF DESCRIPTION OF THE DRAWINGS

FIGS. 1A and 1B show an example of an endoscope for applying an opticalimaging beam and RF radiation to a target in various applications,including using as a bronchoscope bronchial thermoplasty.

FIG. 2 shows an example of an endoscope for applying an optical imagingbeam and RF radiation to a target with a cooling mechanism.

FIG. 3 shows one operation of the endoscope in FIGS. 1A and 1B or inFIG. 2.

FIGS. 4-8 show examples of an optical catheter for directing the opticalimaging beam to the target in the endoscope in FIGS. 1A and 1B or inFIG. 2 and associated optical detection and processing, where theoptical probe head reflects a first portion of an optical probe beamback to the fiber and to direct a section portion of the optical probebeam to a target tissue as the optical imaging beam and further receiveslight returned from the target tissue under illumination of the opticalimaging beam to overlap the light returned from the target tissue withthe first portion to co-propagate in the fiber away from the opticalprobe head.

FIG. 9 shows measurements of water density and water temperature.

FIG. 10 shows measurements of refractive index of water and itstemperature.

DETAILED DESCRIPTION

This application describes examples of techniques, apparatus and systemsthat use optical imaging to achieve temperature profile mapping invarious applications, including a thermotherapy process. One example ofthermotherapy processes is radio frequency (RF) Ablation (RFA) that usesRF energy to destroy malignant cells via thermal treatment and heating.Thermotherapy depends on achieving a proper temperature range to achievedesired cell necrosis efficacy. If a malignancy is being treated usingRFA such as treating liver cancer, RF power levels sufficient to producecomplete cell necrosis must be employed. If the power level is too low,incomplete necrosis will occur increasing the likelihood of recurrence.If power levels are too high, extensive charring of the tissues result,impairing recovery. Mapping of the tissue temperature profile duringthermotherapy is therefore very important to achieve optimum clinicaloutcomes.

Endoscope devices described here use light-based tomographic imaging andtemperature-induced optical property changes to provide temperatureprofile mapping of a target, such as tissue, an organ or other object.

For temperatures above normal body temperature, the mass density ofwater (in grams/cc) decreases as temperature increases. Although thedensity change is not large, sufficiently sensitive instruments andsufficiently sophisticated interpretative techniques can detect andinterpret mass density changes in the water content of soft tissues.When teat is then applied to the tumor to evelvate the local tissuetemperature to approach 42° C., the process of cell death begins. As thetissue temperature increases the time required to assure cell deathdecreases. At about 58° C. proteins begin to denature and at highertemperatures tissue coagulates. The denaturation and coagulation oftissue contained within approximately the same volume, produces densitychanges. As heat is transported from the central coagulation zone andconducted to surrounding tissues, an expanding coagulation zone developsand increases as a function of thermal power delivered over time. Theprimary treatment for both malignant and benign tumors is surgery. Whilesurgery has proven to be effective in cancer treatment, it is expensiveand invasive, frequently requiring lengthy hospital stays for patientrecovery. A critical measure of surgical success is the complete removalof the tumor with surgical margins testing negative to tumor cells uponpathology review.

With the number of cancer cases steadily increasing throughout theworld, a less invasive, less costly way of treating primary andmetastatic tumors is desired. Modern diagnostic imaging modalities andnew interventional methods have set the stage for bringing less invasivemethods to the field of oncology treatment much as imaging has enabledthe diagnosis and minimally invasive treatment of coronary arteries withballoon catheters and drug eluting stents.

The devices and techniques described in this document combines real timeimaging with interventional access to solid tumors offers thepossibility to non surgically treat, or ablate, the tumor mass. Improvedminimally invasive methods are being developed to treat both primary andmetastatic solid tumors. The present devices and techniques can be usedto create thermal injury to tissue when used in conjunction with apercutaneous approach to the lesion. The energy sources used of creatingthe desired thermal injury can vary depending on implementations and mayuse RF (radiofrequency), laser, microwave energy or high intensityfocused ultrasound.

While thermal ablative techniques such as radiofrequency ablation orlaser ablation are rapidly gaining acceptance in the treatment ofinoperable tumors, incomplete treatments are common since there is noreliable method to monitor the treatment zone during the ablation. Atreatment that does not encompass the entire tumor will result inrecurrent growth of the tumor, usually within one or two years. Ablativetechniques will be held back from full adoption as long as the treatmentzone cannot be monitored during the ablation.

The present optical imaging devices and techniques provide opticalmonitoring and feedback control for thermal therapy. Various aspects ofthe such devices and techniques can be found in PCT Application No.PCT/US2005/37730 (PCT Publication No. WO 2006/045013) entitled“Integrated Disease Diagnosis and Treatment System” which isincorporated by reference as part of the disclosure of this document.Such optical imaging systems can be used for mapping of spatiallynon-uniform thermal distributions and can have special utility inthermotherapies where microstructural imaging plays an important role.As an example, bronchial wall microstructure is very important in a new,non-drug asthma thermotherapy called Bronchial Thermoplasty while, invarious CT-guided RFA therapy, low resolution CT imaging is performed ata low rate of CT imaging and thus is inadequate for real time monitoringof RFA therapy. The present optical imaging devices provide fast imagingof the target tissue under treatment, e.g., 30 images/sec, and thus canbe used to measure the thermal effect caused by the RF radiation duringthe RF treatment.

Bronchial Thermoplasty [BT] is a non-drug asthma therapy in clinicaltrials and has been developed over the past seven years as a means ofreducing the severity of asthmatic exacerbations in patients whoseasthma is not well controlled by cortico-steroid or anti-inflammatoryinhalers. BT uses RF energy, applied to the airway smooth muscle [ASM],which surrounds the bronchus in a somewhat random pattern, to reduce theability of the ASM to contract and obstruct airflow. The RF catheter isinserted into the working channel of a bronchoscope and positioned in asecondary/tertiary bronchus, but since the ASM is not visible fromwithin the bronchus, there is no way to position the RF electrode inclose proximity to the ASM. As a consequence, RF energy is appliedperiodically, as the catheter is withdrawn from the bronchus, and, onthe average, damages some of the ASM to some degree. An interestingstatistic is that ASM covers, on the average, only 2.5-4.5% of thesurface of a major airway. The procedure does provide relief, but in alllikelihood, could be more effective if microstructural cross-sectionalimaging allowed the physician to accurately position the RF electrodedirectly above the ASM. In addition, some patients experience adverseevents [including hospitalization] after BT. Since the same treatmentparameters are used for each patient, it seems reasonable to concludethat differences in bronchial wall microanatomy may be responsible forthese adverse events.

The optical imaging systems described in this document can be used toplay the role of a companion diagnostic for BT as well as providingprocedural imaging for guidance and control.

FIGS. 1A and 1B below show an example endoscope apparatus that combinesthe optical imaging for thermal mapping and RFA mechanism. Thisendoscope device includes an optical catheter 4 comprising an opticalfiber 18 to guide an optical imaging beam 14 and an optical probe head1, located at a distal end of the fiber 18 to direct the optical imagingbeam 14 from the fiber 18 to a target and to receive light returned fromthe target under illumination of the optical imaging beam 14. Theoptical probe head 1 as illustrated in this example includes a prismreflector 12 and the distal end facet of the fiber 18 and may beimplemented in other configurations. The optical probe head 1 is fixedto the distal end of the fiber 18. The fiber 18 is attached to a torquecable that can rotate the fiber 18 and thus the optical probe head 1 toallow the optical imaging beam 14 from the optical probe head 1 to scanaround at different positions. In addition, the torque cable can be usedto push or pull the optical probe head 1 to different positions long thelongitudinal direction of the fiber 18. Therefore, under this design,the optical probe head 1 can be used to sense images at differentlocations along the longitudinal direction of the fiber 18 and, at eachlocation, the optical probe head 1 can be rotated to sense thesurrounding areas at different directions.

In this device, a detection module is provided to process the lightreturned from the target under illumination of the optical imaging beam14 to measure a temperature at a location illuminated by the opticalimaging beam 14. An energy source is provided to produce energy that isapplied to the target to raise a temperature at a location of the targetwhere the energy is applied. In this example, the energy source is an RFsource and an RF applicator 6 is provided to direct the RF energy to thetarget. A control mechanism is also provided to control the energysource to set a power level of the energy produced by the energy sourcebased on the measured temperature from the detection module.

An endoscope tube or sheath 2 with a hollow working channel is providedto house the fiber 18 and to hold the RF applicator 6 and an RFtransmission path 8 that connects the RF applicator 6 and the RF source.The RF transmission path 8 can be an electrically conductive tube insome implementations. When implemented as a bronchoscope, the endoscopetube 2 may include other components such as an imaging camera and abiopsy module. The RF applicator 6 can be an RF conductive wires in formof a basket and is attached to an outer sheath 10 outside the opticalfiber 18 in the optical catheter 14 in a way that allows the basket toexpand when the outer sheath 10 is pulled in the endoscope tube orsheath 2. It is a sliding fit so that the optical imaging catheter 4 canrotate within the endoscope sheath 2 to form the image. As illustrated,an applicator mount 7 is provided to engage the RF applicator 6 to theouter sheath 10. The optical catheter 4 also may be moved in alongitudinal manner, to image proximally, distally or exactly at thepoint where the basket struts contact the airway wall tissue. Theconfiguration is such that the coaxial imaging catheter 4 can rotateinside the sheath 10 which is fixed to the basket applicator 6 at thetip. The sheath/imaging catheter unit is initially extended to allow thecollapsed wire RF applicator basket 6 to fit through the working channelorifice as used in bronchoscopes and other endoscopes. The size may notbe so critical in this particular example as the RF BronchialThermoplasty therapy is used in the larger airways.

FIG. 2 shows another example of an endoscope where the optical probehead 1 of the optical catheter 4 includes a prism reflector 12 and afocusing lens 16 that provide optical focusing between the prismreflector 12 and the distal end of the fiber 18. The focusing lens 16 orits interface with the distal end of the fiber 18 can be used by theoptical probe head 1 to reflect a first portion of an optical probe beamguided by the fiber 18 towards the optical probe head back to the fiber18 and to direct a section portion of the optical probe beam to a targettissue as the optical imaging beam 14. The prism reflector 12, thefocusing lens 16 and the distal end of the fiber 18 are fixed inposition relative to each other and rotate or move together as the fiber18 is rotated or moved by a torque cable 20 attached to the fiber 18.The optical probe head 1 receives light returned from the target tissueunder illumination of the optical imaging beam 14 to overlap the lightreturned from the target tissue with the first portion that does notreach the target tissue to co-propagate in the fiber 18 away from theoptical probe head for optical detection and processing.

In many cases, for example in the case of ASM, the therapeutic area ofinterest [ASM] occurs below the surface of a target tissue undertreatment. In this case, if RF power is applied such that thetemperature of the ASM is optimized for cytosis [time/temperature], theairway surface tissue reaches a temperature higher than that of the ASMwhich may lead to damage and an Adverse Event [AE]. Significant severeAEs are reported after BT therapy [10%]. The endoscope design in FIG. 2includes a cooling mechanism that cools the surface of the treatedtarget issue to be lower than a temperature underneath the surface. Inthis regard, a coolant, e.g., a cooling gas or liquid, can be direct topass near the tissue surface to cool the surface. As a specific example,the conductive struts of the RF basket applicator can be made fromflexible hollow tubing which carries the cooling gas or liquid. When aliquid coolant is used, a loop for carrying the liquid coolant is needto direct the coolant out of the endoscope. When a gaseous coolant isused, the gaseous coolant may be directed to the target surface withoutbeing collected. In FIG. 2, two hollow basket struts 22 and 23 are usedto form two cooling loops near the tip of the endoscope which a coolantis passed. An applicator mount 7 is provided to hold the hollow basketstruts 22 and 23 relative to the sheath 10. This design can reduce thesurface temperature of the contact region of the target tissue and thepeak temperature of the treated location on the target issue occurs atsome distance below the surface, depending on the coolant used, and itstemperature and flow rate. This places the peak temperature in closerproximity to the target ASM, resulting in improved therapeutic efficacy.

When the above described endoscope is used as a bronchoscope for BTtreatment, an image can be acquired from each bronchial location, e.g.,over a 1-second imaging time. During BT local tissue temperatures riseand when approaching approximately 40° C. the process of cell deathbegins. As the tissue temperature increases the time required to assurecell death decreases. Between about 55-60° C. proteins begin to losewater and at even higher temperatures tissue chars, producing densitychanges. During the RF power application interval, heat is conducted tosurrounding tissues, expanding the tissue damage zone as a function ofRF energy [RF power×time]. Periodic tomographic images may be taken atfrequent intervals (e.g., few seconds) to monitor process temperatureand/or ASM condition.

After suitable pilot studies have been performed to establish to thecorrelation between adverse events and unique microstructural features[e.g. location/extent of ASM, submucosal position/thickness] it will bepossible to make therapeutic decisions and set therapeutic parametersbased on pre-therapeutic imaging results. This should allow physiciansto utilize optimum RF power settings to achieve efficacy whileminimizing the number and extent of adverse events. Since BT changes thewater content and hence the optical properties of ASM, BT induced ASMdamage can be ascertained by comparing pre- and post-therapymicrostructure images. A growing experience base will allow physiciansto personalize therapeutic parameters to obtain the best clinicaloutcome for the patient. This ability includes locating the ASM as wellas temperature/tissue monitoring during RF power application at eachsite.

The optical catheter 4 in the endoscope in FIGS. 1A and 1B or FIG. 2 canbe used to obtain both (1) imaging of the target tissue and (2) thermalprofile of the target tissue. The imaging information allows thephysician to inspect the images of a treat area and the thermal profileof the target issue allows proper control of the delivery of the RFpower for thermotherapy to achieve effect treatment without damaging thetissue. FIG. 3 shows one exemplary process for operating the endoscopein FIGS. 1A and 1B or FIG. 2. First, an optical imaging beam is directedfrom the endoscope head to the target tissue to obtain reflection of theoptical imaging beam from the target tissue and the therapeutic energyis directed to the target tissue being imaged to administer thermaltherapy (step 31). Next, the light returned by the target underillumination of the optical imaging beam is collected and processed toobtain (1) imaging of the target tissue and (2) thermal profile of thetarget tissue (step 32). The thermal information of the target tissueobtained from the thermal profile of the target tissue is used as afeedback from the target tissue to control an amount of the therapeuticenergy locally applied to different locations of the target tissue tokeep a temperature at each location treated by the therapeutic energybetween a low threshold and a high threshold for the thermalthermotherapy (step 33).

In some implementations, the optical probe head 1 can be used to do animaging scan of the entire treatment area, e.g., performing imaging thesurrounding area at each location of the optical probe head 1 along thelongitudinal direction of the fiber 18 and repeating this imagingoperation at different locations in the longitudinal direction of thefiber 18. This pre-treatment imaging scan allows construction of animaging map of the entire area to be treated and identification ofselected locations in the entire area to be treated. This pre-treatmentimaging scan can be used as a map to control and guide the applicationof the thermotherapy radiation to the selected locations. Duringtreatment at each of the selected locations, the optical probe head 1 isused to monitoring the thermal profile at each selected location and themonitoring result is used to control the power lever of thethermotherapy radiation being applied. Alternatively, in otherimplementations, the optical probe head 1 can be directed to eachlocation of the entire treatment area to perform an imaging scan byrotating the optical probe head 1 to detect whether there is a targetarea to be treated and then perform the radiation treatment if a targetarea is detected. Then the optical probe head 1 is moved along thelongitudinal direction of the fiber 18 to a different location to repeatthe optical imaging for detecting one or more targets and then radiationtreatment of each detected target. During the radiation treatment, theoptical probe head 1 is used to monitoring the thermal profile at eachselected location and the monitoring result is used to control the powerlever of the thermotherapy radiation being applied.

FIGS. 4-8 show examples of an optical catheter 4 for directing theoptical imaging beam 14 to the target in the endoscope in FIGS. 1A and1B or in FIG. 2 and associated optical detection and processing. Theoptical probe head reflects a first portion of an optical probe beamback to the fiber and to direct a section portion of the optical probebeam to a target tissue as the optical imaging beam and further receiveslight returned from the target tissue under illumination of the opticalimaging beam to overlap the light returned from the target tissue withthe first portion to co-propagate in the fiber away from the opticalprobe head. Such design allows for superposition and interplay ofdifferent optical waves and modes propagating along substantially thesame optical path provided by the fiber 18. When one of the opticalwaves or modes interacts with the target, its superposition with anotherwave or mode can be used for acquiring information about the opticalproperties of the substance. This use of a common optical path fordifferent optical waves which may be in the same mode or different modesavoids separation of the reference light beam from the sample light beamin various optical coherence domain reflectometry (OCDR) systems andassociated technical issues caused by the separation of optical pathssuch as uncontrolled fluctuations in the relative optical phase ordifferential delay between the two beams that may adversely affect themeasurements. The use of the common optical path for different opticalwaves in the same or different modes may be advantageously used tostabilize the relative phase among different radiation waves and modesin the presence of environmental fluctuations in the system such asvariations in temperatures, physical movements of the system especiallyof the waveguides, and vibrations and acoustic impacts to the waveguidesand system. In this context, such systems have a “built-in” stability ofthe differential optical path by virtue of their optical designs and arebeneficial for some phase-sensitive measurement, such as thedetermination of the absolute reflection phase and birefringence.

FIG. 4 shows one example of a sensing device according to oneimplementation. This device directs light in two propagation modes alongthe same waveguide to an optical probe head near a sample 205 foracquiring information of optical inhomogeneity in the sample. A sampleholder may be used to support the sample 205 in some applications. Lightradiation from a broadband light source 201 is coupled into the firstdual-mode waveguide 271 to excite two orthogonal propagation modes, 001and 002. A light director 210 is used to direct the two modes to thesecond dual-mode waveguide 272 that is terminated by a probe head 220.The probe head 220 may be configured to perform at least the followingfunctions. The first function of the probe head 220 is to reverse thepropagation direction of a portion of light in the waveguide 272 in themode 001; the second function of the probe head 220 is to reshape anddeliver the remaining portion of the light in mode 002 to the sample205; and the third function of the probe head 220 is to collect thelight reflected from the sample 205 back to the second dual-modewaveguide 272. The back traveling light in both modes 001 and 002 isthen directed by light director 210 to the third waveguide 273 andfurther propagates towards a differential delay modulator 250. Thedifferential delay modulator 250 is capable of varying the relativeoptical path length and optical phase between the two modes 001 and 002.A detection subsystem 260 is used to superpose the two propagation modes001 and 002 to form two new modes, mutually orthogonal, to be receivedby photo-detectors. Each new mode is a mixture of the modes 001 and 002.

The superposition of the two modes 001 and 002 in the detectionsubsystem 260 allows for a range detection. The light entering thedetection subsystem 260 in the mode 002 is reflected by the sample,bearing information about the optical inhomogeneity of the sample 205,while the other mode, 001, bypassing the sample 205 inside probe head220. So long as these two modes 001 and 002 remain independent throughthe waveguides their superposition in the detection subsystem 260 may beused to obtain information about the sample 205 without the separateoptical paths used in some conventional Michelson interferometersystems.

For the simplicity of the analysis, consider a thin slice of the sourcespectrum by assuming that the amplitude of the mode 001 is E₀₀₁ in afirst linear polarization and that of the mode 002 is E₀₀₂ in a second,orthogonal linear polarization in the first waveguide 271. The sample205 can be characterized by an effective reflection coefficient r thatis complex in nature; the differential delay modulator 250 can becharacterized by a pure phase shift Γ exerted on the mode 001. Let usnow superpose the two modes 001 and 002 by projecting them onto a pairof new modes, E_(A) and E_(B), by a relative 45-degree rotation in thevector space. The new modes, E_(A) and E_(B), may be expressed asfollowing:

$\begin{matrix}\left\{ \begin{matrix}{{E_{A} = {\frac{1}{\sqrt{2}}\left( {{^{j\; \Gamma}E_{001}} + {rE}_{002}} \right)}};} \\{E_{B} = {\frac{1}{\sqrt{2}}{\left( {{^{j\; \Gamma}E_{001}} - {rE}_{002}} \right).}}}\end{matrix} \right. & (1)\end{matrix}$

It is assumed that all components in the system, except for the sample205, are lossless. The resultant intensities of the two superposed modesare

$\begin{matrix}\left\{ \begin{matrix}{{I_{A} = {\frac{1}{2}\left\lbrack {E_{001}^{2} + E_{002}^{2} + {{r}E_{001}E_{002}\; {\cos \left( {\Gamma - \phi} \right)}}} \right\rbrack}};} \\{{I_{B} = {\frac{1}{2}\left\lbrack {E_{001}^{2} + E_{002}^{2} - {{r}E_{001}E_{002}\; {\cos \left( {\Gamma - \phi} \right)}}} \right\rbrack}},}\end{matrix} \right. & (2)\end{matrix}$

where φ is the phase delay associated with the reflection from thesample. A convenient way to characterize the reflection coefficient r isto measure the difference of the above two intensities, i.e.

I _(A) −I _(B) =|r|E ₀₀₁ E ₀₀₂ cos(Γ−φ)  (3)

If Γ is modulated by the differential delay modulator 250, the measuredsignal, Eq. (3), is modulated accordingly. For either a periodic or atime-linear variation of Γ, the measured signal responds with a periodicoscillation and its peak-to-peak value is proportional to the absolutevalue of r.

For a broadband light source 201 in FIG. 2, consider the two phases, Γand φ to be dependent on wavelength. If the two modes 001 and 002experience significantly different path lengths when they reach thedetection system 260, the overall phase angle, Γ−φ, should besignificantly wavelength dependant as well. Consequently the measuredsignal, being an integration of Eq. (3) over the source spectrum, yieldsa smooth function even though Γ is being varied. The condition for asignificant oscillation to occur in the measured signal is when the twomodes 001 and 002 experience similar path lengths at the location oftheir superposition. In this case the overall phase angle, Γ−φ, becomeswavelength independent or nearly wavelength independent. In other words,for a given relative path length set by the modulator 250, anoscillation in the measured signal indicates a reflection, in the othermode, from a distance that equalizes the optical path lengths traveledby the two modes 001 and 002. Therefore the system depicted in FIG. 2can be utilized for ranging reflection sources.

Due to the stability of the relative phase between the two modes, 001and 002, phase-sensitive measurements can be performed with the systemin FIG. 2 with relative ease. The following describes an exemplarymethod based on the system in FIG. 2 for the determination of theabsolute phase associated with the radiation reflected from the sample205.

In this method, a sinusoidal modulation is applied to the differentialphase by the differential delay modulator 250, with a modulationmagnitude of M and a modulation frequency of Ω. The difference inintensity of the two new modes is the measured and can be expressed asfollows:

I _(A) −I _(B) =|r|E ₀₀₁ E ₀₀₂ cos [M sin(Ωt)−φ].  (4)

It is clear from Eq. (4) that the measured exhibits an oscillation at abase frequency of Ω and oscillations at harmonic frequencies of the basefrequency Ω. The amplitudes of the base frequency and each of theharmonics are related to φ and |r|. The relationships between r and theharmonics can be derived. For instance, the amplitude of thebase-frequency oscillation and the second harmonic can be found from Eq.(4) to be:

A _(Ω) =E ₀₀₁ E ₀₀₂ J ₁(M)|r|sin φ;  (5 a)

A _(2Ω) =E ₀₀₁ E ₀₀₂ J ₂(M)|r|cos φ,  (5b)

where J₁ and J₂ are Bessel functions of the first and second order,respectively. Eq. (5a) and (5b) can be used to solve for |r| and φ, i.e.the complete characterization of r. We can therefore completelycharacterize the complex reflection coefficient r by analyzing theharmonic content of various orders in the measured signal. Inparticular, the presence of the base-frequency component in the measuredis due to the presence of φ.

FIG. 5 shows an exemplary implementation of the system depicted in FIG.4. The spectrum of source 201 may be chosen to satisfy the desiredranging resolution. The broader the spectrum is the better the rangingresolution. Various light sources may be used as the source 201. Forexample, some semiconductor superluminescent light emitting diodes(SLED) and amplified spontaneous emission (ASE) sources may possess theappropriate spectral properties for the purpose. In this particularexample, a polarization controller 302 may be used to control the stateof polarization in order to proportion the magnitudes of the two modes,001 and 002, in the input waveguide 371. The waveguide 371 and otherwaveguides 372 and 373 may be dual-mode waveguides and are capable ofsupporting two independent polarization modes which are mutuallyorthogonal. One kind of practical and commercially available waveguideis the polarization maintaining (PM) optical fiber. A polarizationmaintaining fiber can carry two independent polarization modes, namely,the s-wave polarized along its slow axis and the p-wave polarized alongits fast axis. In good quality polarization maintaining fibers these twomodes can have virtually no energy exchange, or coupling, forsubstantial distances. Polarization preserving circulator 310 directsthe flow of optical waves according to the following scheme: the twoincoming polarization modes from fiber 371 are directed into the fiber372; the two incoming polarization modes from fiber 372 are directed tothe fiber 373. A polarization-preserving circulator 310 may be used tomaintain the separation of the two independent polarization modes. Forinstance, the s-wave in the fiber 371 should be directed to the fiber372 as s-wave or p-wave only. Certain commercially availablepolarization-preserving circulators are adequate for the purpose.

The system in FIG. 5 implements an optical probe head 320 coupled to thewaveguide 372 for optically probing the sample 205. The probe head 320delivers a portion of light received from the waveguide 372, the lightin one mode (e.g., 002) of the two modes 001 and 002, to the sample 205and collects reflected and back-scattered light in the same mode 002from the sample 205. The returned light in the mode 002 collected fromthe sample 205 carries information of the sample 205 and is processed toextract the information of the sample 205. The light in the other mode001 in the waveguide 372 propagating towards the probe head 320 isreflected back by the probe head 320. Both the returned light in themode 002 and the reflected light in the mode 001 are directed back bythe probe head 320 into the waveguide 372 and to the differential delaymodulator 250 and the detection system 260 through the circulator 310and the waveguide 373.

In the illustrated implementation, the probe head 320 includes a lenssystem 321 and a polarization-selective reflector (PSR) 322. The lenssystem 321 is to concentrate the light energy into a small area,facilitating spatially resolved studies of the sample in a lateraldirection. The polarization-selective reflector 322 reflects the mode001 back and transmits the mode 002. Hence, the light in the mode 002transmits through the probe head 320 to impinge on the sample 205. Backreflected or scattered the light from the sample 205 is collected by thelens system 321 to propagate towards the circulator 310 along with thelight in the mode 001 reflected by PSR 322 in the waveguide 372.

In the above described examples, the optical probe head sends out lightin two different propagation modes where light in one of the two modescarries the information from the sample. Alternatively, light in asingle propagation mode may be used as the input light to the opticalprobe head and as output light from the optical probe head. Hence,devices based on this design not only use a common optical path todirect light to and from the probe head and sample but also control thelight in a single mode. In comparison with above examples where twodifferent modes are used for light coming out of the probe heads, thissingle-mode design further eliminates or reduces any differences betweendifferent modes that propagate in the same optical path.

FIG. 6 shows one exemplary system for acquiring information of opticalinhomogeneity and other properties in substances with only onepropagation mode inside waveguides. A broadband or low-coherence lightfrom Broadband Light Source 201 is directed to a probe head 2110 bymeans of polarization-maintaining waveguides 271 and 272. A partialreflector inside the probe head 2110 reverses the direction of a smallportion of the input light to create a radiation wave 1 whiletransmitting the remainder of the input light to the sample 205.Backscattered or reflected light from the sample 205 becomes a secondradiation wave 2 and is collected by the probe head 2110. The probe head2110 combines and couples both the radiation waves 1 and 2 back into thewaveguide 272. The radiation waves 1 and 2 travel in the waveguide 272towards Light the light director 210 which directs radiation waves 1 and2 through the waveguide 273 towards the detection module 2101. Notably,the radiation waves 1 and 2 output from the probe head 2110 are in thesame mode as the input light to the probe head 2110. the probe head 2110does not change the mode of light when directing the radiation waves 1and 2 to the waveguide 272.

The detection module 2101 includes a beam Splitter 2120, two opticalpaths 2121 and 2122, an optical variable delay element 2123 in the path2122, a beam combiner 2130, and two optical detectors 2141 and 2142. Thebeam splitter 2120 splits the light in the waveguide 273, which includesthe radiation waves 1 and 2 in the same mode, into two parts thatrespectively propagate in the two optical paths 2121 and 2122. Notably,each of the two parts includes light from both the radiation waves 1 and2. The variable delay element or delay line 2123 in the optical path2122 is controlled by a control signal to adjust the relative opticaldelay between the two optical paths 2121 and 2122 and may be implementedby, e.g., the exemplary delay elements described in this application andother delay designs. The beam combiner 2130 combines the signals of thetwo optical paths to overlap with each other and to output two opticalsignals for optical detectors 2141 and 2142, respectively. The beamcombiner may be a polarization beam splitter which splits the combinedlight into two parts, orthogonal in polarization to one another.

The probe head 2110 may include a partial reflector to produce theradiation wave 1 which does not reach the sample 205. Assuming thesingle propagation mode for the light to the probe head 2110 and thelight out of the probe head 21110 is a polarization mode, the lightreflected from the partial reflector in the probe head 2110, i.e., theradiation wave 1, has the same polarization as the light collected fromthe sample, the radiation wave 2. Therefore, both Radiation 1 and 2travel in the same propagation mode in the waveguides, 272 and 273.Because the radiation waves 1 and 2 are reflected from differentlocations, they experience different optical path lengths when reachingthe beam splitter 2120. The effect of variable delay element 2123 is toadd an adjustable amount of the delay in the light in the path 2122relative to the light in the path 2121.

In operation, the variable delay element 2123 can be adjusted so thatthe partial radiation 1 reaching the polarization beam splitter 2130through the path 2122 can be made to experience a similar optical pathlength as the partial radiation 2 reaching the beam splitter 2130 viathe other path 2121. The superposition of the two beams at the photodetectors 2141 and 2142 causes a measurable intensity variation as theirrelative path length is being varied by the variable delay element 2123.This variation can be utilized to retrieve information on theinhomogeneity and other properties of the sample 205.

FIG. 7 shows an exemplary implementation of the system in FIG. 6 usingpolarization maintaining optical fibers. A polarization controller 202may be placed at the output of the light source 201 to control thepolarization of the input light in one polarization mode. The opticalhead 2110 is shown to include a lens system 2111 and a partial reflector2112. Two mirrors 1 and 2 are used to construct the two optical pathsbetween the beam splitters 2120 and 2130. The optical radiationreflected from the partial reflector 2122 and from the sample 205 travelin the polarization-maintaining (PM) fiber 272 in the same mode. Themain portions of the radiation waves 1 and 2 are deflected to the mirror1 while the remaining portions are directed to the mirror 2 by the beamsplitter 2120.

The incident plane of the polarizing beam splitter 2130 can be made tohave a finite angle with respect to the polarization directions of lightfrom both the Mirror 2 in one optical path and the variable delayelement 2123 from the other optical path. In this configuration, lightenergies received by both detectors 2141 and 2142 are the superpositionof the two radiations, i.e., Radiation 1 and Radiation 2. It should beappreciated that the linkage between the beam splitters 2120 and 2130can be made by means of optical fibers or other optical waveguides toeliminate the free space paths and the two mirrors 1 and 2.

In some implementations, the probe head may be designed to cause a firstportion of the first mode to reverse its propagation direction whiledirecting the remaining portion, or a second portion, to reach thesample. The reflection or back scattered light of the second portionfrom the sample is collected by the probe head and is controlled in thesecond propagation mode different from the first mode to produce areflected second portion. Both the reflected first portion in the firstpropagation mode and the reflected second portion in the secondpropagation mode are directed by the probe head through a commonwaveguide into the detection module for processing. In comparison withthe implementations that use light in two modes throughout the system,this alternative design further improves the stability of the relativephase delay between the two modes at the detection module and providesadditional implementation benefits.

FIG. 8 illustrate one exemplary design of the optical layout of theoptical sensing system and its system implementation with an electroniccontroller. An input waveguide 871 is provided to direct light in afirst propagation mode, e.g., the mode 001, from the broadband lightsource 201 to a light director 810. The waveguide 871 may be a modemaintaining waveguide designed to support at least one propagation modesuch as the mode 001 or 002. When light is coupled into the waveguide871 in a particular mode such as the mode 001, the waveguide 871essentially maintains the light in the mode 001. A polarizationmaintaining fiber supporting two orthogonal linear polarization modes,for example, may be used as the waveguide 871. Dual-mode waveguides 272and 273 are used to direct the light. A light director 510 is used tocouple the waveguides 871, 272, and 273, to convey the mode 001 from theinput waveguide 871 to one of the two modes (e.g., modes 001 and 002)supported by the dual-mode waveguide 272, and to direct light in twomodes from the waveguide 272 to the dual-mode waveguide 273. In theexample illustrated in FIG. 8, the light director 810 couples the lightin the mode 001 from the waveguide 871 into the same mode 001 in thewaveguide 272. Alternatively, the light director 810 may couple thelight in the mode 001 from the waveguide 871 into the different mode 002in the waveguide 272. The dual-mode waveguide 271 is terminated at theother end by a probe head 820 which couples a portion of light to thesample 205 for sensing.

The probe head 820 is designed differently from the prove head 320 inthat the probe head 830 converts part of light in the mode 001 into theother different mode 002 when the light is reflected or scattered backfrom the sample 205. Alternatively, if the light in the waveguide 272that is coupled from the waveguide 871 is in the mode 002, the probehead 820 converts that part of light in the mode 002 into the otherdifferent mode 001 when the light is reflected or scattered back fromthe sample 205. In the illustrated example, the probe head 820 performsthese functions: a) to reverse the propagation direction of a smallportion of the incoming radiation in mode 001; b) to reshape theremaining radiation and transmit it to the sample 205; and c) to convertthe radiation reflected from the sample 205 to an independent mode 002supported by the dual-mode waveguide 272. Since the probe head 820 onlyconverts part of the light into the other mode supported by thewaveguide 272, the probe head 820 is a partial mode converter in thisregard. Due to the operations of the probe head 820, there are two modespropagating away from the probe head 820, the mode 001 that bypasses thesample 205 and the mode 002 for light that originates from samplereflection or back scattering.

In the examples shown in FIGS. 4-8, an optical delay is introduced atthe output of the common optical waveguide to perform a time-domainimaging in extracting images from the returned light by the opticalprobe head. The depth scan is obtained based on a Fourier-transform fromthe measured signals at different phase delays at different times.Alternatively, the same designs for the optical probe head and thecommon optical waveguide can be used to extract images from the returnedlight by the optical probe head based on optical frequency domainimaging (OFDI) under a swept source optical coherence tomographyconfiguration. Under OFDI, a wavelength-swept light source is used toproduce a probe beam to the optical probe head for probing theamplitude, phase, polarization and spectral properties of backscattering light from a target, e.g., a tissue. The depth scan isobtained based on a Fourier-transform from the acquired spectra. Assuch, the optical delay used in the time-domain technique can beeliminated. OFDI offers intrinsic signal-to-noise ratio (SNR) advantageover the time domain techniques because the interference signal can beeffectively integrated through a Fourier transform enabling significantimprovements in imaging speed, sensitivity and ranging depth requiredfor in-vivo tissue imaging.

The optical imaging information obtained from the devices in FIGS. 4-8based on either the time-domain imaging and frequency-domain imaging canbe processed to map the thermal field or profile of a target. Changes indensity, due to temperature changes, affect the optical properties [suchas refractive index] of water and tissue. FIGS. 9 and 10 showmeasurements in water to illustrate the correlations between thetemperature, density and the refractive index. FIG. 10 shows therefractive index of water as a function of temperature at standarddensity and 589 nm wavelength. A linear fit shows that the curve can beparameterized by n=1.341−2.262×10⁻⁵T.

In some implementations, the thermal treat radiation exposure is notapplied to a tissue occurs during a monitoring period when the opticalimaging beam is directed to the target to measure the target. In otherimplementations, both the optical imaging beam and the thermal treatmentradiation can be applied to the target at the same time. The result ofan optical index change can cause a phase shift which can lead to a fullfringe for a small [1 deg. C.] temperature change allowing a twodimensional representation of the tissue density change to be assembled.The phase changes on each scanning of the relative phase between theoptical imaging beam and the reflected beam that does not reach thetarget are translated into contrast changes. Scanning the opticalimaging beam in a circumferential manner, while rapidly performing aseries of imaging scans allows a raster scan image to be constructedwhere the thermally-induced density changes show up as contrastvariations. The contrast variations reflect the combined effect oftemperature changes and tissue denaturing [water loss], two importantparameters for BT. By comparing the pre- and intra-BT images, thetemperature increase due to thermotherapy can be extracted.

Since cytotoxicity requires only a few minutes exposure to temperaturesbetween 50-60 deg C. and denaturation requires temperatures in excess of60 deg C., thermally-induced image contrast is almost completely due tothe change in temperature rather than water loss. However, above 60 degC., water loss [denaturation] becomes inceasingly important as anadditional contribution to image contrast. The post-BT image [aftercooling] contrast will reflect changes in ASM density due to BT.Notably, the optical techniques described in this document can be usedto obtain imaging of the, microstructure, spectral properties andtemperature profile of a target tissue, organ or object. Consequently,if we obtain a tomographic tissue image and apply an energy source[light, RF] so as to subsequently generate a temperature change[non-uniform in many cases], we can map that temperature gradient by theimage contrast changes it produces. The overall functions of the presenttechniques include 1) monitoring therapeutic temperature [‘processcontrol’] to maintain it in the correct ‘thermotherapeutic window’ and2) pre- and post-tissue imaging to assess the extent of tissue ‘damage’as a result of thermotherapy.

While this specification contains many specifics, these should not beconstrued as limitations on the scope of an invention that is claimed orof what may be claimed, but rather as descriptions of features specificto particular embodiments. Certain features that are described in thisspecification in the context of separate embodiments can also beimplemented in combination in a single embodiment. Conversely, variousfeatures that are described in the context of a single embodiment canalso be implemented in multiple embodiments separately or in anysuitable sub-combination. Moreover, although features may be describedabove as acting in certain combinations and even initially claimed assuch, one or more features from a claimed combination can in some casesbe excised from the combination, and the claimed combination may bedirected to a sub-combination or a variation of a sub-combination

Only a few examples and implementations are described. One of ordinaryskill in the art can readily recognize that variations, modificationsand enhancements to the described examples may be made.

1. An endoscope device, comprising: an optical catheter comprising an optical fiber to guide an optical imaging beam and an optical probe head, located at a distal end of the fiber to direct the optical imaging beam from the fiber to a target and to receive light returned from the target under illumination of the optical imaging beam; a detection module to process the light returned from the target under illumination of the optical imaging beam to measure a temperature at a location illuminated by the optical imaging beam; an energy source that produces energy that is applied to the target to raise a temperature at a location of the target where the energy is applied; and a control mechanism that controls the energy source to set a power level of the energy produced by the energy source based on the measured temperature from the detection module.
 2. The device of claim 1, wherein: the energy source is an RF source, and the device comprises an RF applicator engaged to the sheath to receive RF energy from the RF source and to apply the received RF energy to the target.
 3. The device of claim 2, wherein: the RF applicator includes a hollow tube that both conducts the RF energy for application of the RF energy to the target and carries a coolant inside the hollow tube to cool a surface of the target to allow a temperature at a location underneath the surface to be higher than a temperature of the surface under application of the RF energy.
 4. The device of claim 2, wherein: the RF applicator surrounds the optical catheter and allows the optical catheter to rotate within the sheath.
 5. The device of claim 4, comprising: a housing comprising a hollow working channel, in which the optical catheter is placed to rotate within the hollow working channel, and an RF transmission path connecting the RF applicator and the RF source RF applicator and located inside the hollow working channel.
 6. The device of claim 5, wherein: the RF applicator comprises an expandable basket that can carry RF energy, the expandable basket structured to expand when the expandable basket is located outside a distal opening of the hollow working channel and to collapse when the expandable basket is pulled into the distal opening of the hollow working channel.
 7. The device of claim 1, wherein: the optical probe head of the optical catheter comprises a prism that reflects the optical imaging beam from the fiber to a location of the target and collects the light returned from the target into the fiber.
 8. The device of claim 1, wherein: the optical probe head of the optical catheter splits the optical imaging beam from the fiber into a first portion that is directed to the target and a second portion that does not reach the target and is reflected back to the fiber to co-propagate with the light returned from the target that is collected by the optical probe head along the fiber, and the detection module receives both the second portion and the light returned from the target via the fiber and processes the received second portion and the light returned from the target to extract the measured temperature.
 9. The device as in claim 8, wherein: the detection module comprises an optical differential delay unit to produce and control a relative phase delay between the second portion and the light returned from the target via the fiber and performs a time-domain Fourier transform of the obtained signals at different relative phase delays.
 10. The device as in claim 8, wherein: the detection module performs a frequency-domain Fourier transform of the spectra of the light from the optical probe head.
 11. The device of claim 1, comprising: a beam scanning mechanism to scan the optical imaging beam on the target to direct the optical imaging beam to different positions on the target, wherein the detection module changes a phase delay between two portions of received light of the second portion and the light returned from the target via the fiber at each location of the optical imaging beam on the target to obtain an image of the each location of the target.
 12. The device of claim 1, wherein: the detection module processes received light of the second portion and the light returned from the target via the fiber to measure a change of a density of the target caused by application of the energy from the energy source.
 13. The device of claim 1, comprising: a mechanism to cool a surface of the target to allow a temperature at a location underneath the surface to be higher than a temperature of the surface under application of the energy from the energy source.
 14. The device of claim 13, wherein: the mechanism comprises a tube that carries a cooling flow to pass a location near the distal end of the optical catheter.
 15. The device of claim 1, wherein: the energy source produces a laser beam, separate from the optical imaging beam, as the energy applied to the target.
 16. The device of claim 1, wherein: the energy source generates a microwave energy signal as the energy applied to the target.
 17. The device of claim 1, wherein: the energy source generates an ultrasound signal as the energy applied to the target.
 18. A method for guiding the application of a thermotherapeutic radiation to a target tissue, comprising: using an optical catheter comprising an optical fiber to guide an optical imaging beam to direct the optical imaging beam from the fiber to a target and to receive light returned from the target under illumination of the optical imaging beam; processing the light returned from the target tissue under illumination of the optical imaging beam to measure a temperature at a location illuminated by the optical imaging beam; applying a thermotherapeutic energy to the target tissue to raise a temperature at a location of the target tissue where the energy is applied; and controlling an amount of the thermotherapeutic energy applied to the target tissue, based on the measured temperature by using the optical imaging beam, to control the temperature at the target tissue between a low limit above which a thermotherapeutic effect is present and a high limit above which a damage to the target tissue occurs.
 19. The method as in claim 18, comprising: using the optical imaging beam to obtain an image of a location of the target tissue illuminated by the optical imaging beam, in addition to obtaining the measured temperature; and using both the measured temperature and the obtained image from the optical imaging beam to monitor and control application of the thermotherapeutic energy to the target tissue.
 20. The method as in claim 18, comprising: using the optical catheter to split the optical imaging beam from the fiber into a first portion that is directed to the target and a second portion that does not reach the target and is reflected back to the fiber to co-propagate with the light returned from the target tissue.
 21. The method as in claim 20, comprising: producing a relative phase delay between the second portion and the light returned from the target tissue via the fiber to cause an optical interference between the second portion and the light returned from the target tissue; detecting a change to the optical interference caused by the application of the thermotherapeutic energy to the target tissue to measure a change in density of the target tissue; and using the measured change in density of the target tissue to obtain the measured temperature.
 22. The method as in claim 20, comprising: performing a frequency-domain Fourier transform of spectra of light in the fiber that combines the light returned from the target tissue and the second portion that does not reach the target to extract a change to the optical interference caused by the application of the thermotherapeutic energy to the target tissue and to measure a change in density of the target tissue; and using the measured change in density of the target tissue to obtain the measured temperature.
 23. The method as in claim 18, comprising: scanning the optical imaging beam on the target tissue to obtain a two-dimensional map of temperatures at locations of the target tissue; and using the two-dimensional map of temperatures at locations of the target tissue to control the application of the thermotherapeutic energy to the target tissue.
 24. The method as in claim 18, wherein: the processing of the light returned from the target tissue under illumination of the optical imaging beam comprises: obtaining image contrasts from the light returned from the target tissue before and after the thermotherapeutic energy; comparing the obtained image contrasts from the light returned from the target tissue before and after the thermotherapeutic energy to obtain a contrast variation caused by application of the thermotherapeutic energy; and using the contrast variation to obtained the measured temperature.
 25. The method as in claim 18, comprising: cooling a surface of the target tissue to allow a temperature at a location underneath the surface to be higher than a temperature of the surface under application of the thermotherapeutic energy.
 26. The method as in claim 25, wherein: using a tube that is located near the distal end of the optical catheter to carry a cooling flow to providing the cooling of the surface of the target tissue.
 27. An endoscope device for providing guided thermotherapy, comprising: an endoscope tube comprising a hollow working channel; an optical catheter comprising an optical fiber located inside the hollow working channel to guide an optical probe beam and, an optical probe head, located at a distal end of the fiber, to reflect a first portion of the optical probe beam back to the fiber and to direct a section portion of the optical probe beam to a target tissue as an optical imaging beam, the optical probe head receiving light returned from the target tissue under illumination of the optical imaging beam to overlap the light returned from the target tissue with the first portion to co-propagate in the fiber away from the optical probe head; an optical delay device coupled to the fiber to receive the first portion and the light returned from the target tissue to produce a variable relative phase delay between the first portion and the light returned from the target tissue; an optical detector that detects the light of the first portion and the light returned from the target tissue from the optical delay device; a processing unit to receive output from the optical detector and to extract a temperature at a location illuminated by the optical imaging beam from information of the target carried by the light returned from the target tissue; an RF applicator engaged to the endoscope tubing and near the optical probe head to apply RF energy to the target tissue to raise a temperature at a location of the target where the RF energy is applied; and a control mechanism that controls an amount of the RF energy to be applied by the RF applicator to the target tissue based on the measured temperature.
 28. The device as in claim 27, wherein: the RF applicator includes a hollow tube that both conducts the RF energy for application of the RF energy to the target tissue and carries a cooling liquid or gas inside the hollow tube to cool a surface of the target to allow a temperature at a location underneath the surface to be higher than a temperature of the surface under application of the RF energy.
 29. The device as in claim 27, wherein: the optical probe head includes a partial reflector that reflects the first portion of the optical probe beam back to the fiber and transmits the second portion of the optical probe beam as the optical imaging beam to the target tissue.
 30. The device as in claim 27, wherein: the optical delay device includes a beam splitter to separate the light from the fiber into a first light beam along a first optical path and a second light beam along a second optical path; a variable optical delay element in one of the first and the second optical paths to cause the relative phase delay between the first light beam and the second light beam; and a beam combiner to combine the first light beam and the second light beam to produce combined light. 